Method and device for detecting cellular targets in bodily sources using carbon nanotube thin film

ABSTRACT

A device and method detect cellular targets in a bodily source by utilizing a biofunctional pad comprised of a thin film of carbon nanotubes (CNT&#39;s). When antibodies are absorbed by the CNT&#39;s, cellular targets having markers matching the antibodies may be detected in a bodily source placed upon the biofunctional pad by measuring the conductivity of the thin film using conductive contacts electrically coupled to the thin film, as the binding of the receptors in the cellular targets to the antibodies changes the free energy in the thin film. In many respects, the device functions as a Field Effect Transistor (FET) with the bodily source, e.g., blood, acting as a polyelectrolyte liquid gate electrode to create a varying electrostatic charge or capacitance in the thin film based upon the binding of cellular targets in the source to the antibodies present on the biofunctional pad.

CROSS-REFERENCE TO RELATED APPLICATIONS

This application claims priority to U.S. Provisional Application Ser.No. 61/312,913 filed by Balaji Panchapakesan on Mar. 11, 2010, andentitled “METHOD AND DEVICE FOR DETECTING CELLULAR TARGETS IN BODILYSOURCES USING CARBON NANOTUBE THIN FILM,” which application isincorporated by reference in its entirety.

GOVERNMENT RIGHTS

This invention was made with Government support under Grant No. ECCS0853066 awarded by the National Science Foundation. The Government hascertain rights in this invention.

FIELD OF THE INVENTION

The invention is generally related to detecting cellular targets inbodily sources, e.g., detecting circulating cancer cells in blood. Theinvention is also generally related to the use of carbon nanotubes(CNT's) in medical diagnostic applications.

BACKGROUND OF THE INVENTION

Identification and quantitation of numerous biological molecules togenerate a complex molecular profile is required for diagnosis,monitoring, and prognostic evaluation of complex diseases such ascancer. Despite outstanding progress in the area of cancer biology,significant challenges remain in translating biological knowledge ofcancer surface markers into clinically relevant devices that could beused as diagnostic or monitoring tools for cancer management. Developinghigh-throughput and low cost diagnostic cell and tissue analysis fordisease detection has remained a challenge.

For example, breast cancer is the most diagnosed cancer in women, and ithas been found that developing breast cancers shed transformed cellsinto the blood, with more malignant breast cancer cells appearing in theblood in later stages. It is believed by many that early detection ofcirculating breast cancer cells might improve diagnosis of early breastcancer and ultimately reduce breast cancer-related deaths. Therefore,significant efforts have been made toward the development of methods anddevices for detecting circulating breast cancer cells in blood.

Circulating tumor cells (CTC's) have long been analyzed ex vivo by flowcytometry and fluorescence microscopy to measure characteristic cellsurface markers, such as epithelial cell adhesion marker (EpCAM), ageneral purpose epithelial cell marker that is common to circulatingtumor cells. Many of these techniques, however, are expensive and timeconsuming, often requiring several days to generate results.

More recently, it has been found that small bundles of single wallcarbon nanotubes (SWCNT), ˜10 nm diameter, lithographically patternedbetween two electrodes, with adsorbed monoclonal antibodies, willdisplay a sensitivity to a single cancer cell in 1 μL of blood.Moreover, such devices have the potential to detect the presence ofcancer cells in blood in a matter of minutes, rather than days as isoften the case with other methodologies. However, the use of single orsmall bundles of SWCNT's presents challenges in a clinical setting dueto the difficulty in fabricating such single or small bundle SWCNTsamples, and ensuring that the cancer cells are bridging the electrodesto achieve reliable detection. Furthermore, the use of nanoscale devicesprecludes the use of large blood volumes that are typically analyzed ina clinical setting.

Therefore, a need continues to exist in the art for an improvedmethodology and device for detecting cancer cells and other cellulartargets in blood and other bodily sources.

SUMMARY OF THE INVENTION

The invention addresses these and other problems associated with theprior art by providing a device and method of detecting cellular targetsin a bodily source utilizing a biofunctional pad comprised of a thinfilm of carbon nanotubes (CNT's). When antibodies are absorbed by theCNT's, cellular targets having markers matching the antibodies may bedetected in a bodily source placed upon the biofunctional pad bymeasuring the conductivity of the thin film using conductive contactselectrically coupled to the thin film, as the binding of the receptorsin the cellular targets to the antibodies changes the free energy in thethin film. In many respects, the device functions as a Field EffectTransistor (FET) with the bodily source, e.g., blood, acting as apolyelectrolyte liquid gate electrode to create a varying electrostaticcharge or capacitance in the thin film based upon the binding ofcellular targets in the source to the antibodies present on thebiofunctional pad.

Consistent with one aspect of the invention, a device for detectingcellular targets in a bodily source includes a substrate; abiofunctional pad comprising a thin film of carbon nanotubes (CNT's)disposed on the substrate and adapted to receive antibodies associatedwith a cellular target; and a plurality of conductive contacts disposedon the substrate and electrically coupled to the thin film. Theplurality of conductive contacts are configured for use in detecting thecellular target in a bodily source by measuring a conductivity of thethin film when the antibodies are received by the thin film and thebodily source is disposed on the biofunctional pad and in contact withthe antibodies, whereby the conductivity of the thin film is indicativeof the presence of the cellular target in the bodily source.

Consistent with another aspect of the invention, a method of fabricatinga sensor for detecting cellular targets in a bodily source includesforming a thin film of carbon nanotubes (CNT's) on a carrier usingvacuum filtration; mechanically bonding the thin film to a dielectriclayer on a semiconductor substrate; separating the thin film from thecarrier; patterning the thin film to form a biofunctional pad; anddepositing a plurality of conductive contacts on the substrate, with atleast a portion of each conductive contact overlapping and electricallycoupled to the thin film.

Consistent with yet another aspect of the invention, a method ofdetecting cellular targets in a bodily source includes placing a bodilysource on a biofunctional pad comprising a thin film of carbon nanotubes(CNT's) upon which is disposed antibodies associated with a cellulartarget; and measuring the conductivity of the thin film using aplurality of conductive contacts electrically coupled to the thin film,whereby the conductivity of the thin film is indicative of the presenceof the cellular target in the bodily source.

These and other advantages and features, which characterize theinvention, are set forth in the claims annexed hereto and forming afurther part hereof. However, for a better understanding of theinvention, and of the advantages and objectives attained through itsuse, reference should be made to the Drawings, and to the accompanyingdescriptive matter, in which there is described exemplary embodiments ofthe invention.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a top plan view of a cellular detection and profiling sensorarray consistent with the invention.

FIG. 2 is a top plan view of one of the sensors from the sensor array ofFIG. 1.

FIG. 3 is a cross-sectional view of the sensor of FIG. 2, taken alonglines 3-3.

FIG. 4 is a flowchart illustrating exemplary steps in fabricating thesensor array of FIG. 1.

FIGS. 5A-5H are cross-sectional views illustrating the fabrication ofthe sensor array of FIG. 1 during various of the steps illustrated inFIG. 4.

FIG. 6 is a top plan view of an alternate sensor to that of FIG. 2,incorporating separate drive and sensing contacts.

FIG. 7 is a flowchart illustrating exemplary steps in detecting andprofiling cellular targets using the sensor array of FIG. 1.

FIG. 8 is a chart illustrating I-V characteristics of an exemplarysensor consistent with the invention functionalized with IGF1Rantibodies for blood samples with varying numbers of MCF7 cells.

FIG. 9 is a chart illustrating a change of resistance in an exemplarysensor consistent with the invention as a number of MCF7 cells in bloodsamples when functionalized with several different antibodies.

FIG. 10 is a chart illustrating overexpression ratios plotted as afunction of antibody type, illustrating relatively high overexpressionin IGF1R and EpCAM antibodies, medium overexpression in Her2 and EGFRantibodies, and low overexpression in IgG and PSMA antibodies.

FIGS. 11A and 11B are scanning electronic microscope (SEM) image of aCNT thin film surface, uncoated (FIG. 11A) and coated with 10 μL of 5μg/mL antibodies.

FIG. 12A is a representative I vs. Vds plots of experimental stages ofan anti-Her2 functionalized CNT thin film transistor.

FIG. 12B is a plot of I vs. time comparisons for 1 μL of anti-IGF1R, 1μL of pure blood, and 1 μL of blood spiked with 5 MCF-7 cells.

FIG. 13 is a plot of normalized electrical signals as a function of thenumber of MCF-7 cells for A) non-specific IgG antibody B) non-specificPSMA antibody C) specific IGF1R antibody D) specific HER2 antibody andE) specific EpCAM antibody functionalized CNT films.

FIG. 14 illustrates changes in a normalized electrical signal fordevices functionalized with IgG, IGF1R, Her2, EpCAM, and EGFRantibodies.

FIGS. 15A and 15B are respective plots of I-Vg for anti-IgG and MCF7cells interaction (FIG. 15A), and of I-Vg for anti-IGF1R and MCF7 cellsinteraction (FIG. 15B).

FIG. 16 is a bar graph illustrating percentage changes in an electricalsignal applied across a CNT thin film for samples of samples of rubygold nanoparticles, isolated white blood cells, isolated MCF7 cancercells, blood and blood mixed with MCF7 cells.

FIGS. 17A and 17B are plots of current vs. time for blood samplesdisposed on CNT thin films with hydrophobic (FIG. 17A) and hydrophilic(FIG. 17B) surfaces.

DETAILED DESCRIPTION

Embodiments consistent with the invention use thin films of carbonnanotubes (CNT's) for detecting surface receptors or markers in cellulartargets in bodily sources, e.g., cancer cells in blood. As will becomemore apparent below, when blood mixed with cancer cells is brought intocontact with a thin film of CNT's that has been functionalized withmonoclonal antibodies, the conductivity of the thin film changes, andtypically does so in a manner that is directly related to the number ofcancer cells in blood. Therefore, by applying a known voltage across thethin film, e.g., through a pair of conductive contacts or electrodeselectrically coupled to the thin film, the presence of cancer cells canbe determined from the current sensed through the film, with the currentdecreasing a function of the number of cells in the blood.

Among other benefits, the techniques described herein provide devicesthat are readily adaptable to clinical environments, and may haveapplicability in third world countries or in other instances whereaccess to health care facilities is limited. Thin films of CNT's arereadily adaptable to batch fabrication techniques and CMOS/MEMSfabrication techniques. In addition, as compared to conventionaltechnologies, surface markers can be detected in blood in a few minutesvs. a few days, and the level of skill required of the technician may besubstantially reduced.

One application of the invention, for breast cancer detection in blood,will hereinafter be the focus of the instant application. It will beappreciated by one of ordinary skill in the art that the invention mayhave applicability in connection with the detection of other forms ofcancer, e.g., prostate cancer, or the detection of other cellulartargets. In addition, the invention may have applicability in connectionwith detecting cellular targets in other bodily sources, e.g., otherbodily fluids.

Turning now to the Drawings, wherein like numbers denote like partsthroughout the several views, FIG. 1 illustrates a device 10 consistentwith the principles of the invention. Device 10 includes a sensor arrayof sensors 12 disposed on a substrate, e.g., a silicon or othersemiconductor wafer 14. In the illustrated embodiment, each sensor 12may be separately configured with different antibodies to targetdifferent cellular targets, and it may be desirable to provide multiplesensors 12 with the same antibodies, e.g., to provide the ability todouble check results, or to test blood samples from different patientson the same device. While device 10 is illustrated with a 3×4 array ofsensors 12, it will be appreciated that any number of sensors 12 may bedisposed in a given device consistent with the invention. For example,it may be desirable in some embodiments to utilize 200 sensors 12 sothat a 1 mL blood sample may be analyzed in 54 drops applied to the 200sensors.

FIG. 2 illustrates one of sensors 12 in greater detail, while FIG. 3illustrates a cross-section of one of sensors 12, taken along lines 3-3of FIG. 2. Sensor 12 in the illustrated embodiment is formed on top of adielectric layer 16 on substrate 16, e.g., a silicon dioxide layer.Sensor 12 includes a biofunctional pad 18 formed from a thin film ofCNT's and a pair of contacts or electrodes 20 through which theconductivity of the thin film may be measured. Contacts 20 may be formedof gold or another conductive material, and partially overlap the thinfilm in regions 22.

In use, biofunctional pad 18 is functionalized to detect a particularmarker on a cellular target by applying antibodies 24 on a surfacethereof. Then, a drop of blood or other bodily source, illustrated at26, is deposited on biofunctional pad 18, separated from contacts 20,and a voltage is applied across contacts 20 to generate a current thatis measured to calculate the conductivity, i.e., the IV characteristics,of the thin film forming the biofunctional pad.

As noted above, each sensor 12 effectively functions as a Field EffectTransistor (FET) with the bodily source, e.g., blood, acting as apolyelectrolyte liquid gate electrode to create a varying electrostaticcharge or capacitance in the thin film based upon the binding ofcellular targets in the source to the antibodies present on thebiofunctional pad.

As discussed in the aforementioned paper, CNT's are generally p-typematerials, and as a result, applying a positive gate voltage to the thinfilm of CNT's depletes the carriers and reduces the overall conductancethrough the thin film. The dependence of conductance on gate voltage isideal for biosensing applications, as the binding of charged species tothe gate dielectric is analogous to applying a voltage through a gateelectrode. Thus the conductance of a p-type CNT would decrease when aprotein with a positive surface charge binds to an antibody. It has beenfound that blood spiked with cancer cells decreases or increases theconductance of the sensor with increasing number of cells depending onthe net charge. This is a general pattern for many antibodies, althoughother antibodies, e.g., EGFR, may increase the conductance of the thinfilm with increasing number of cells. The mechanism is therefore one ofelectrostatic gating of the CNT thin film. It is believed that bloodspiked with cancer cells acts as a gate electrode. Varying positive(negative) voltage at the gate electrode decreases (increases) theconductance of the device. In this case, increasing the number of cellsis equivalent to increasing the voltage of the liquid gate. While thisseems simplistic, one can also look at the capacitance of the liquidgate as a function of the Debye length to understand the reason forexcellent gate coupling of blood with increase in cancer cells.

The total gate capacitance, which determines the charging of the CNT'sunder a certain gate voltage, consists of electrostatic (Ce) and quantum(Cq) components. For back-gating devices, the capacitance of the gate isgiven by C_(bg)=2π∈∈₀/ln (2 h/r), where (∈∈₀) is the gate materialdielectric constant (e.g., 3.9×8.85×10⁻¹² F/m), h is the gate oxidethickness (e.g., 500 nm), and r is the thickness of the CNT film (e.g.,180 nm). The calculated capacitance per unit length of 500 nm silicondioxide back gating is about 1.245×10⁻¹⁰ F/m. For blood as a top gateone can approximate it as a liquid electrolyte top gate with acapacitance given by: C_(liquid gate)=2π∈∈₀/ln (r+λ_(D)/r), where λ_(D)is the Debye length or the electronic screening length resulting fromions. Now if one assigns a value of ˜10 nm for Debye length, thecapacitance per unit length of the liquid gate is about 6.445×10⁻⁸ F/m.Due to the higher capacitance, better gate channel coupling is achieved.Further, it can be seen that blood as a gate has a capacitance of twoorders better than a conventional back gated structure. Although theDebye length varies with different salt concentrations, the estimatedDebye length used in the calculation is still valid because the totalcapacitance variation caused by the change in the Debye length is smallas long as it is still on the order of a few nanometers. Cell surfacereceptors with a net positive or negative charged increase or decreasethe current in blood depending on their surface charge. Thus, byoptimizing a sensor with thinner CNT films, one can achieve excellentgating in a liquid environment that can be used as a mechanism forsensing circulating cancer cells.

In addition, it is believed that change in current in a sensorconsistent with the invention is also related to the extracellular andintra-cellular potentials of the cellular targets. This makes the sensorand method of its use highly specific for specific cell types.

It should also be noted that, after etching and patterning the thin filmof CNT's, the surface of the thin film becomes highly hydrophobic due tothe nature of the single walled CNT's (SWCNT's). The hydrophobic natureof the surface causes a deposited blood droplet to remain on a specificspot on the biofunctional pad without shorting the contacts. Thehydrophobic nature of the surface also causes the antibodies to diffuseslowly and arrange themselves on the surface of the CNT's, and makes itpossible for the adsorbed antibodies to interact with the cell surfaceantigens and create the change in conductivity.

With additional reference to FIGS. 5A-5H, a sensor consistent with theinvention may be fabricated using a process 50 shown in FIG. 4. Process50 begins in block 52 by forming a CNT wafer, e.g., by a vacuumfiltration technique such as disclosed in Lu et al., “Nanotubemicro-optomechanical actuators,” Applied Physics Letters 88, 253107(2006). The wafer includes a thin film 18 of CNT's, e.g., about 100 nmto about 150 nm in thickness, deposited on a mixed cellulose ester (MCE)carrier or filter 28 (FIG. 5A).

Next, in block 54, the substrate is prepared by forming a dielectriclayer 16 on a silicon or other semiconductor wafer 14 (FIG. 5B), e.g.,through oxidation of the silicon wafer to create a layer of silicondioxide.

Next, in block 56, the CNT wafer is bonded to the substrate throughmechanical compression while heating to about 75 degree Celsius, bondingthin film 18 to dielectric layer 16 (FIG. 5C). Thereafter, the CNT waferfilter 28 is removed in block 58 using an acetone vapor bath to dissolvethe filter away from the thin film of CNT's (FIG. 5D).

Next, in block 60, the CNT thin film is etched by patterning aphotoresist mask 30 using a lithographic process to cover the regions ofeach biofunctional pad, and then etching the remaining CNT thin filmusing an etching technique such as deep reactive-ion etching (DRIE)(FIG. 5E). Thereafter, the photoresist mask is removed.

Next, in block 62, electrodes or contacts are deposited, first bypatterning a photoresist mask 32 using a lithographic process to exposethe regions of each contact (FIG. 5F) and then depositing a conductivematerial such as gold, aluminum, copper, platinum, or other conductivemetal or alloy providing a low contact resistance, e.g., via sputteringor other suitable deposition technique. The contacts may be deposited,for example, to a thickness of about 100 nm. Thereafter, the photoresistmask is removed, resulting in contacts 20 being formed with overlappingregions 22 (FIG. 5G).

Next, in block 64, the assembly is annealed, e.g., at about 150 to about200 degrees Celsius in an Argon or other inert gas atmosphere for about20 minutes. Doing so improves the contact between each contact orelectrode and the thin film, thereby lowering the contact resistance sothat the bulk of the contact resistance in the sensor is due to thebinding of cellular targets to the antibodies absorbed into the thinfilm of CNT's.

Next, in block 66, it may be desirable to test the sensors on the wafer,e.g., by measuring the IV characteristics of the sensors by applying avoltage across the contacts 20. In addition, Raman spectroscopy may beperformed to characterize the CNT's in the thin film of each sensor andidentify any potentially defective films.

Subsequent to testing, it may also be desirable during fabrication, asshown in block 68, to deposit antibodies 24 on the biofunctional pad(FIG. 5H). Antibodies may be deposited, for example, via drop coating ofpure antibodies, or via covalent bonding or other known techniques. Inone suitable technique, antibodies may be drop coated onto abiofunctional pad and allowed to set for about 10 minutes to enable theantibodies to diffuse and be absorbed into the thin film, and then washaway the remaining liquid using deionized (DI) water. Depending upon thetype of antibodies and the environmental robustness thereof, it may alsobe desirable to autoclave or freeze the sensors during the fabricationprocess to preserve the antibodies on the surface of the biofunctionalpad until the sensors are ready to be used.

In an alternate embodiment, however, the sensors may be fabricatedwithout antibodies deposited thereon, requiring the antibodies to bedeposited immediately prior to use in a clinical environment.

In the embodiment illustrated in FIGS. 1-5H, each sensor includes abiofunctional pad 18 of about 1.5 mm×1.5 mm, with each contact 20 beingabout 1 mm×1 mm and overlapping the biofunctional pad in regions 22,sized about 0.1 mm×0.3 mm. In this embodiment, it may be desirable todrop coat about 5 μL to about 10 μL of antibodies at a concentration ofabout 5 μg/mL, which results in an about 20 nm thick layer of antibodiesdeposited on the biofunctional pad.

While a pair of contacts, disposed over two adjacent corners of thebiofunctional pad, is used in each sensor 12, other configurations andnumbers of contacts may be used in a sensor consistent with theinvention. For example, FIG. 6 illustrates an alternate sensor 80including a thin layer of CNT's 82 and four contacts 84, 86 overlappingin regions 88. One pair of contacts 84 disposed on opposite corners ofbiofunctional pad 82 may be used as drive pads, through which a currentis passed, and the other pair of contacts 86 may be used as sensingpads, through which the resistance or conductivity of the biofunctionalpad is measured. In other embodiments, contacts may be disposed in otherpositions, e.g., overlapping the edges, of a biofunctional pad.

FIG. 7 next illustrates a process 100 for testing a blood sample usingdevice 10 of FIG. 1. Process 100 begins in block 102 by optionallydepositing the antibodies on the biofunctional pads in the mannerdiscussed above, if not already so done during fabrication. Next, inblock 104, blood droplets (e.g., about 5 μL) are deposited on thebiofunctional pads, without contacting the contacts, either manually orvia a robotic system. Then, in block 106, the IV characteristics of thesensor are recorded over time by applying a known voltage, e.g., up toabout 25 mV across the contacts thereof and measuring the current. Inmany instances, each sensor will stabilize within several minutes, e.g.,5 minutes or so, once all cellular targets in the blood bind with theantibodies. Thereafter, in block 108, the results may be analyzed todetermine what cellular targets were found in the blood based upon whatmarkers were expressed with the different antibodies on differentsensors in the array. In addition, it may also be possible to predict ordetermine the number of cellular targets within each blood sample, e.g.,the number of cancer cells predicted in a given drop of blood, as the IVcharacteristics of each sensor will change based upon the number ofcells, and thus the number of bindings that occur with the antibodies.

The types of antibodies used to test a given blood sample may bedifferent in different embodiments and applications. For example, totest for the presence of breast cancer, it may be desirable to utilizeIGF1R, Her2, EpCAM, and EGFR antibodies on different sensors, while totest for the presence of prostate cancer, it may be desirable to utilizePSMA antibodies. It may even be desirable to utilize antibodies thatexpress for different types of cancer on the same sensor array so that asingle blood sample may be tested for the presence of multiple types ofcancer. Other combinations of antibodies may also be used for otherdiagnostic applications of the invention.

The provisional application cross-referenced herein discusses testresults performed with sensors fabricated in the manner disclosedherein. A portion of these results are illustrated in FIGS. 8-10.Detection and profiling of 10-300 MCF7 breast cancer cells in 5 μLaliquots of blood (the typical reported range for circulating tumorcells in the blood of patients with metastatic breast cancer) wasperformed. Incubation of blood spiked with cancer cells resulted inunambiguous decreases in sensor microjunction conductance, where pureblood resulted in higher conductance (lower resistance) of themicrojunctions compared to blood spiked with cancer cells. The sensorwas able to detect a minimum of 10 MCF7 cells in 5 μL of blood, as wellas the maximum number of 300 MCF7 cells in blood for several differentantibodies. FIG. 8, for example, illustrates the I-V characteristics ofsensors functionalized with IGF1R antibodies using blood with 10-300MCF7 cells, and shows a measurable decrease in conductivity with anincrease in the number of MCF7 breast cancer cells in the blood.

FIG. 9 illustrates the change in resistance and conductance of sensorsvs. the number of MCF7 cancer cells in a 5 μL blood sample for variousantibodies, including IGF1R, EpCAM, and Her2, as well as non-specificIgG and PSMA antibodies, and shows that anti-IGF1R, anti-EpCAM, andanti-Her2 showed a measurable conductance change for MCF7 cells in bloodas compared to the non-specific antibodies.

One potential way to scale cellular measurements is to determine acalibration curve between the cellular overexpression of a surfaceantigen and the change in electrical signal. In current clinicalpractice, diagnoses are mainly reported as the presence or absence ofmalignant cells in the specimen. The capability to quantify, profile,and stratify cancer cells would likely improve diagnosis. A criticalissue when screening cancer cells is how to correlate the expressionlevels of tumor markers to the number of malignant cells in a givensample. Without this knowledge one could either measure high expressionin relatively few cells or low expression in many cells.

One can define an overexpression ratio asΔR^(Max# of cells)/ΔR^(Min# of cells). This description is quiteappropriate as this ratio increases with overexpression. One of theoutcomes of results illustrated in FIG. 9 is that when one replots thedata as overexpression ratios, one finds that the ratio of ΔR³⁰⁰/ΔR¹⁰(IGF1R)=7.07, ΔR³⁰⁰/ΔR¹⁰ (EpCAM)=5.4, ΔR³⁰⁰/ΔR¹⁰ (Her2)=3.9, ΔR³⁰⁰/ΔR¹⁰(EGFR)=2.9, ΔR³⁰⁰/ΔR¹⁰ (IgG)=1.27, and ΔR³⁰⁰/ΔR¹⁰ (PSMA)=0.77. This isshown in FIG. 10.

As can be seen in this figure, while the change in resistance of IGF1Rin FIG. 9 was lower than EpCAM and Her2, the overexpression ratios werethe highest for IGF1R. This shows that the definition of overexpressionis indeed valid, and is in fact consistent with Western Blot analysis,which shows a similar overexpression of IGF1R in MCF7 cells compared toHer2. Further, the results also indicate specific numbers for EpCAM,EGFR and Her2 which are all valid surface markers for breast cancer.These overexpression ratios from 1.0 to 7.0 may also be assignedmalignancy. A ratio of 1.0 may be considered benign or negative for thatmarker and 7.0 may be considered malignant or positive. Furthermore,based on these numbers one can scale the number of cells. From theaforementioned results, it is believed that plotting the number of cellsagainst their overexpression ratios may give a linear change that canactually predict the number of cells in blood. Furthermore, by usingmore markers, one may be able to increase the accuracy of thistechnique.

It has also been found that, in some embodiments, it may also bedesirable to alter the hydrophobicity or hydrophilicity of abiofunctional pad to alter the response characteristics of thebiofunctional pad. It has been found, in particular, that a CNT thinfilm is typically hydrophobic in nature, and that the presence of acellular target in a bodily source disposed on a biofunctional pad tendsto decrease the conductivity of the CNT thin film in the biofunctionalpad such that the conductivity of the thin film is inverselyproportional to the presence of the cellular target in the bodilysource. However, by treating the biofunctional pad to alter the physicalstructure of the CNT thin film, the hydrophilicity of the CNT thin filmmay be increased, and notably, the response of the biofunctional pad maybe altered such that conductivity increases, rather than decreases, inresponse to the presence of a cellular target in a bodily sourcedisposed on the biofunctional pad, such that the conductivity of thethin film is proportional to the presence of the cellular target in thebodily source.

FIG. 11A, for example, illustrates an SEM image of the entangled natureof CNT's in a thin film. This type of entangled network presents idealsurface characteristics for cells to stick to such a surface. It hasbeen found that the entangled nature of CNT's presents ideal surfacesfor antibodies to stick to the surface even for non-covalentfunctionalization methods. The CNT surfaces show high degree ofhydrophobicity due to the exposure of carbon nanotubes in a oxygenplasma during device patterning thereby creating rough surface.Adsorption of an antibody can decrease the surface energy therebydecoupling its surface wettability from bulk properties and enablinghydrophobicity. In addition, in some embodiments, the surface chemistrycan be tailored with molecules such as silane to even create asuperhydrophobic surface with high contact angles. Such surfaces areself-cleaning and therefore can enable variety of medical relateddevices.

In one experimental implementation, for example, the electricalresponses (I vs. Vds) of CNT thin film devices functionalized withspecific antibodies were recorded in order to determine if differentelectrical signals were produced by antibody, blood, or blood with MCF-7cells. FIG. 11B, for example, illustrates an SEM image of a CNT thinfilm functionalized with 10 μL of 5 μg/mL antibodies.

Upon adding biological components, noticeable changes in conductancewere observed, as shown in a typical electrical measurement for ananti-HER2 coated device in FIG. 12A. The reduction in conductance wasnegligible for a phosphate buffered saline (PBS) wash. However, a 50%drop in the current of the device was observed after the adsorption of 5μL of anti-HER2. After antibody adsorption, the addition of blood mixedwith cancer cells resulted in an additional ˜30% decrease in deviceconductance. This result is further observed from a real time currentmeasurement (I vs. T), shown in FIG. 12B. Current decreased ˜10% afteradding antibodies, ˜25% for a blood control sample and ˜60% for bloodmixed with cancer cells.

The electrical behaviors of CNT thin films functionalized with specific(anti-IGF1R, anti-HER2 and anti-EpCAM) and non-specific (anti-IgG) ornon-cognate (anti-PSMA) antibodies were measured in order to determinewhether specific detection of MCF-7 cells was possible in a sample ofunaltered blood. For an initial study, 5 μL of anti-IGF1R or anti-IgGwere immobilized on the surface of the CNT networks followed by theaddition of blood samples with a ramp of MCF-7 cell concentrations. FIG.13 shows that devices printed with IgG experienced less than a ˜10%change in conductivity while devices printed with IGF1R exhibited a ˜60%drop in conductivity with increasing number of MCF-7 breast cancer cellsin blood.

A summary of these specificity studies are presented in FIG. 13. Here itcan be observed that for specific antibodies such as anti-IGF1R,anti-HER2 and anti-EpCAM, the electrical signal (resistance changesbetween current baseline and after adding MCF-7 cells with blood)increased as a function of increasing number of MCF-7 cells in bloodsamples. However, the same is not true for non-specific IgG andnon-cognate PSMA. The electrical signatures remained the same despitethe addition of blood mixed with MCF-7 cells. The specific interactionbetween antibodies and receptors on the cell surface may be defined inthe form of a ratio called the overexpression ratio. The overexpressionratio is the ratio of change in electrical signal for the maximum numberof cells spiked in blood (300 cells) to the change in electrical signalfor the minimum number of cells in blood (10 cells). This relates to thespecificity of the sensor. The ratio is highest for IGF1R (˜7.0), EpCAM(˜6.0), Her2 (˜3.6) and IgG (˜0.8). The ratios give some degree ofspecificity based on the binding of the antibodies to the receptors incells. The number of binding sites for EpCAM and Her2 surface markersand their ratio have been shown in 9 different cancer cell lines usingstandard titration methods. For MCF7 cells the EpCAM expression wasreported as 222.1 (713.7)×103 binding sites while Her2 expression was25.2 (71.6)×103 binding sites respectively. Comparing these measurementsto the overexpression ratios, the EpCAM over expression ratio was higher(˜6.0) than the ratio for Her2 (˜3.6) for MCF7 cells in blood, whichsuggests that the CNT-antibody array data gives similar results tostandard titration methods for the number of binding events. In otherwords, it can be inferred that the change in electrical signal arisesfrom the number of cooperative binding events happening on the surfaceof the device.

In another experimental implementation, I-V plots were recorded on 5 μLstabilized blood samples from three patients with metastatic breastcancer, the results of which are shown in Table I below:

TABLE 1 Molecular analysis of metastatic breast cancer patients VeridexCTC, CNT-mAb Patient ER PR Her2 in 7.5 mL CTC, in 5 μl^(†) A + + − ND*~50 B + + − 97 ~50 C − − + 1 ~300 *deceased before Veridex available^(†)based on overexpression ratio

In this implementation, the CNT thin film devices were spotted withantibodies against IGF1R, HER2, EpCAM, EGFR, and nonspecific IgG, andtested with patient blood samples. The change in conductivity was almost10-fold greater for specific antibodies over non-specific IgG.

Electrical measurements equivalent to those performed from controlledblood samples were recorded for the three metastatic breast cancerpatients. When I-V characteristics of patient C blood interacting withall the different antibodies was compared with immunohistochemicalanalysis done on Patient C, it was found that the cells wereHer2-positive in both cases. Additionally, Veridex CellSearch analysiswas performed for the patients. It was observed from both Patient B andC in Table 1 that when the cells were Her2-positive (Patient C, 1 celldetected), the Veridex gave low cell numbers compared to Her2-negative(Patient B, 97 cells detected). It is believed that the low cell numbersassociated with Her2 status may indicate that EpCAM targeting alonecannot capture all CTC's in a blood sample. The surfaces of CTC's areheterogenous and therefore many different types of markers may benecessary for accurate capture, profiling, and enumeration of CTC's.FIG. 14 illustrates the changes in a normalized electrical signal forPatient B for devices functionalized with IgG, IGF1R, Her2, EpCAM, andEGFR antibodies.

In another experimental implementation, the surface interactionsoccurring in devices based on CNT thin film transistors were studiedusing liquid gated CNT FET's. The goal of these experiments was toidentify the interactions between the CNT's and the antibody-receptorbinding that lead to the charge carrier depletion or decrease incurrent. The transfer characteristics (I vs. Vg) of liquid gatedtransistors were monitored upon the addition of 1 μL of antibodies and 1μL of 5 MCF-7 cells mixed with blood in order to identify theelectrostatic interactions taking place between CNT's and the binding ofsurface receptors with antibodies. For these experiments, only onesingle concentration of MCF-7 cells (5 MCF-7 cells/μL) and blood wasused. The transfer characteristics were recorded for devicesfunctionalized with nonspecific IgG and anti-IGF1R antibodies as shownin FIGS. 15A and 15B (the anti-IgG and anti-IGF1R plots are above theMCF-7 plots in these figures). Distinct differences between theelectrical characteristics of non-specific and specific interactions ofMCF-7 surface receptors were observed. For a device printed with IGF1Rantibodies, there was a shift in the threshold voltage (˜250 mV),whereas devices printed with anti-IgG showed no distinguishable shiftalthough conductance was reduced for negative gate voltage. The currentdecreased for both non-specific and specific antibodies in the negativegate voltage region. However, there was a shift in the gate voltage forthe specific antibody in the positive side, suggesting that geometricdeformations occur around the cellular interactions giving rise to astress, leading to scattering sites on a CNT, and thus to reducedconductance. At the same time the device characteristic is modified onlyfor negative gate voltages, leaving the transconductance in the positivegate voltage region unaffected.

For this latter experiment, an alternate device implementation may beused, where a localized liquid gate configuration modulates current in aconducting channel. The CNT FET's were scaled down to a smaller filmarea of ˜0.008 mm² with only ˜10-100 μm gap between patterned electrodesin order to observe any charge transfer and minimize the diffusivebehavior of charged particles. Electrical currents were measured forspecific and non-specific antibody-cells surface marker interactions. Inaddition, the relationship between electric current as a function ofnumber of CTC's in a sample was explored. The change in signal level wasrelated to the overexpression of targeted cell surface antigens.

As noted above, it may also be desirable in some embodiments to alterthe surface of a CNT thin film to change the hydrophobicity orhydrophilicity of the surface. In one exemplary embodiment, for example,annealing may be performed, e.g., as discussed above in connection withblock 64 of FIG. 4, but at a higher temperature than described inconnection with this figure. It is believed that annealing at a highertemperature, e.g., about 200 to about 400 degrees Celsius, or about 300degrees Celsius or higher, increases the hydrophilicity of a CNT thinfilm surface by “burning” the CNT's on the surface of the thin film,causing the CNT's to curl or curve, and effectively reducing the densityof CNT's at the surface of the film. It is believed that by doing so,the continuity of the surface is interrupted, exposing holes or pits inthe surface that receive liquid and thus increase the hydrophilicity ofthe surface. In one exemplary implementation, for example, it was foundthat a CNT thin film transforms from hydrophobic to hydrophilic at about300 degrees Celsius.

It will be appreciated that altering the surface of a CNT thin film maybe performed in a number of manners consistent with the invention. Inaddition to annealing a wafer after deposition of the thin film andelectrodes, annealing may be performed at other points in thefabrication process, e.g. prior to electrode deposition. In addition,heat may be applied to the thin film in other manners, e.g., viainfrared heating, etc. In addition, other surface treatments may beperformed, including, for example, chemical treatment, oxygen plasmatreatment, etc. In general, any treatment that lowers the density ofCNT's on a thin film surface and increases hydrophilicity may be usedconsistent with the invention.

It has been found that altering the surface of a CNT thin film to renderthe surface hydrophilic causes the conductivity of a biofunctional padfunctionalized with an antibody to increase in response to the presenceof a biological target, which is opposite to the response of a CNT thinfilm with a hydrophobic surface. As such, in implementations where it isdesirable to utilize a positive conductivity relationship with thepresence of a biological target, treating the surface of the CNT thinfilm may be desirable.

In addition, the ability to selectively alter the surface of only someof the biofunctional pads on a wafer, e.g., as might be performed usinginfrared heating, provides the ability to provide both hydrophilic andhydrophobic biofunctional pads on the same wafer. In some embodiments,for example, the outputs of hydrophilic and hydrophobic biofunctionalpads functionalized with the same antibody may be combined to increasethe sensitivity and/or signal to noise ratio of a sensor, or to providea reconfigurable functionalized surface for a sensor.

In still other embodiments, different surfaces may be combined withdifferent antibodies to provide a more thorough analysis of the types ofbiological targets present in a bodily source. Different types of CTC's,for example, may exhibit different responses to different antibodies, sothat not only the presence of a CTC in a blood sample, but the type ofCTC, may be detected through the analysis of the conductivity ofdifferent biofunctional pads functionalized with different antibodies.

In an additional experimental implementation, a device comprising a CNTthin film with a hydrophobic surface was fabricated spanning between apair of gold electrodes deposited on a glass substrate. 5 μL samples ofruby gold or gold nanoparticles of 3-5 nm were tested along with 5 μLsamples of isolated white blood cells, isolated MCF7 cancer cells, bloodand blood mixed with MCF7 cells, and the change in an electrical signalapplied across the electrodes was measured for each sample. As shown inFIG. 16, the gold nanoparticles showed the smallest percentage change,followed by isolated white blood cells. Isolated MCF7 cells exhibited alarger percentage change; however, blood and blood mixed with MCF7 cellsexhibited unique conductivity on the CNT surface, with the blood mixedwith MCF7 cells exhibiting the highest conductivity.

In addition, as shown in FIG. 17A, a measurement of current over timefor the aforementioned device illustrates the decrease in conductivityover time seen in a hydrophobic CNT thin film upon which 5 μL of bloodhas been adsorbed. FIG. 17B, in contrast, illustrates a measurement ofcurrent over time for a similar device for which the CNT thin film istreated to render the surface hydrophilic, and with 5 μL of bloodadsorbed thereon. In contrast with the hydrophobic surface, the plot ofcurrent vs. time for the hydrophilic surface exhibits an increase inconductivity over time. In each of FIGS. 17A and 17B, the plots of twoblood samples are shown but are not normalized to one another, as thepurpose of these figures is merely to illustrate the relative changes inconductivity that occur over time with CNT films having hydrophobic andhydrophilic surfaces.

Various additional modifications beyond those discussed herein will beapparent to one of ordinary skill in the art. Therefore, the inventionlies in the claims hereinafter appended.

What is claimed is:
 1. A device for detecting cellular targets in abodily source, comprising: a substrate; at least one biofunctional padcomprising a thin film of carbon nanotubes (CNT's) disposed on thesubstrate and adapted to receive antibodies; a plurality of conductivecontacts disposed on the substrate, the conductive contacts electricallycoupled to the thin film; a layer of antibodies associated with acellular target applied directly to a surface of the thin film of thebiofunctional pad; at least a portion of the surface of the thin filmbeing altered with a process in the formation of the device to providethe thin film with a hydrophilic surface so that with a layer ofantibodies applied thereto, the conductivity of the biofunctional pad isproportional to the presence of the cellular target on the biofunctionalpad; the plurality of conductive contacts configured for use inmeasuring a change in conductivity of the biofunctional pad when thebodily source is disposed on the biofunctional pad and in contact withthe layer of antibodies for detecting the presence of the cellulartarget in the bodily source.
 2. The device of claim 1, wherein the layerof antibodies is at least partially absorbed by the thin film.
 3. Thedevice of claim 1, wherein the cellular target comprises a cancer cell,and wherein the bodily source comprises a drop of blood.
 4. The deviceof claim 1, wherein the thin film is rectangular in shape, and whereinthe plurality of conductive contacts comprises first and second contactsdisposed at and partially overlapping first and second corners of thethin film.
 5. The device of claim 1, wherein the layer of antibodies isselected to target cellular receptors from the group consisting ofIGF1R, Her2, EpCAM, and EGFR, and wherein the cellular target comprisesa breast cancer cell.
 6. The device of claim 1, wherein the thin filmand the plurality of conductive contacts define a first sensor, thedevice further comprising a plurality of sensors disposed on thesubstrate, wherein each of the plurality of sensors respectivelyincludes a thin film of CNT's and a plurality of conductive contactselectrically coupled to the respective thin film, and wherein theplurality of sensors are each adapted to receive a layer of antibodiesselected from among a plurality of antibody types to detect differentmarkers potentially associated with a cellular target.
 7. The device ofclaim 1, wherein the bodily source functions as a polyelectric liquidgate for a transistor defined by the thin film and the plurality ofcontacts.
 8. The device of claim 1, wherein at least another portion ofthe surface of a thin film of a biofunctional pad is hydrophobic, andwherein the conductivity of the another portion of a thin film isinversely proportional to the presence of the cellular target on theanother portion of the surface of a thin film of the biofunctional pad.9. The device of claim 1, wherein the portion of the surface is alteredwith a process that includes annealing the substrate to provide thinfilm with the hydrophilic surface after the plurality of conductivecontacts is disposed thereon.
 10. The device of claim 1 wherein at leasta portion of the antibodies applied directly to a surface of the thinfilm are absorbed by the thin film of CNT's.
 11. The device of claim 1,wherein the portion of the surface of the thin film is altered with aprocess that includes annealing the substrate to provide thin film withthe hydrophilic surface after forming the thin film of CNT's.
 12. Thedevice of claim 11, wherein the substrate is annealed at a temperatureof about 200 degrees Celsius to about 400 degrees Celsius.
 13. Thedevice of claim 11, wherein the substrate is annealed at a temperatureof at least about 300 degrees Celsius.
 14. The device of claim 1,wherein the portion of the surface of the thin film is altered with aprocess that includes the application of heat to the thin film toprovide thin film with the hydrophilic surface.
 15. The device of claim1 wherein the portion of the surface of the thin film is altered with aprocess that includes at least one of chemical treatment or oxygenplasma treatment or infrared heating.
 16. The device of claim 1 whereinthe portion of the surface of the thin film is altered with a processthat lowers the density of the carbon nanotubes on the thin filmsurface.
 17. A device for detecting cellular targets in a bodily source,comprising: a substrate; at least one biofunctional pad comprising athin film of carbon nanotubes (CNT's) disposed on the substrate andadapted to receive antibodies; a plurality of conductive contactsdisposed on the substrate, the conductive contacts electrically coupledto the thin film; a layer of antibodies associated with a cellulartarget applied directly to a surface of the thin film of thebiofunctional pad; at least a portion of the surface of the thin filmbeing altered with a process in the formation of the device to providethe thin film with a hydrophilic surface and at least another portion ofthe surface of the thin film not being altered with the process so that,with a layer of antibodies applied to the biofunctional pad, theconductivities of the respective portions of the surface of the thinfilm are differently proportional with respect to each other in thepresence of the cellular target on the biofunctional pad; the pluralityof conductive contacts configured for use in measuring changes inconductivity of the biofunctional pad when the bodily source is disposedon the biofunctional pad and in contact with the layer of antibodies fordetecting the presence of the cellular target in the bodily source. 18.The device of claim 17, wherein the portion of the surface is alteredwith a process that includes annealing the substrate to provide the thinfilm with the hydrophilic surface after the plurality of conductivecontacts is disposed thereon.
 19. The device of claim 17, wherein theportion of the surface of the thin film is altered with a process thatincludes annealing the substrate to provide the thin film with thehydrophilic surface after forming the thin film of CNT's.
 20. The deviceof claim 17 wherein the portion of the surface of the thin film isaltered with a process that includes at least one of chemical treatmentor oxygen plasma treatment or infrared heating.